Optical imaging, such as fluorescence or optoacoustic imaging, can be used for in vivo imaging of molecular functions and gene expression in live biological tissues. For example, using externally administered agents with sensitivity and specificity to certain functional, molecular, and/or cellular targets (such as fluorochromes or nanoparticles), optical imaging can be used to visualize events that are not detectable using conventional imaging modalities, such as ultrasound or X-ray. Furthermore, the combination of these agents with an appropriate optical detection system can lead to very high detection sensitivity and high biological specificity. As a result, optical imaging approaches are becoming increasingly important for the diagnosis and monitoring of disease.
With respect to fluorescence imaging, generally, excitation light is transmitted toward a tissue to excite the emission of fluorescent light from fluorochromes associated with the tissue. This method is conventionally performed in fluorescence microscopy for high-resolution imaging of histological sections of biological tissue. In addition to this conventional use, examples of in vivo fluorescence imaging approaches include confocal imaging, multiphoton imaging, and total internal reflection fluorescence microscopy. The excitation light often used in these approaches is in the near infrared (near-IR) range, as higher tissue penetration depths can be achieved in comparison to those when light in the visible wavelength range is used. However, even using near-IR light, the in-tissue penetration limit of light during the fluorescence imaging is less than about 0.5 millimeters. As a result, in their current implementation, the fluorescence imaging systems are not appropriate for three-dimensional or quantitative imaging of hollow organs, such as in intra-vascular, pulmonary/bronchoscopic, or gastro-intestinal imaging.
For example, near-IR fluorescence catheter systems have been developed for detecting distributions of fluorescence in tissues, including imaging of hollow organs such as the gastrointestinal tract, pulmonary system, and cardiovascular system. In their present form, such systems rely predominantly on surface information derived from fluorescence reflectance imaging, which provides a number of drawbacks. In particular, the fluorescence signal emanating from target fluorescent probes embedded in the wall of a hollow organ suffers from attenuation due to scattering and absorption in tissue and blood. This attenuation is generally exponentially dependent on the unknown distance of the probes from the catheter when the organ is filled with blood. Thus, fluorescence light emanating from untargeted probes that are closer to the catheter than the actual targeted probes may overshadow the true signal and lead to inaccurate quantification. An example of such a scenario is provided by a situation when the hollow organ is a blood vessel and a fluorescent dye is circulated through the blood stream. In this case, the entire fluorescent image may be saturated from the fluorescence signal within the blood and, therefore, will not indicate whether any fluorescent probe exists deeper in the blood vessel wall. Because of these limitations, fluorescence catheters, endoscopic systems, and angioscopic systems substantially lack the ability to provide quantitative three-dimensional or even two-dimensional information. This type of information may be critical in some cases to accurately map disease, quantify response to therapies, and/or geographically localize fluorescence signals within target pathology.
An alternative to fluorescence imaging is multi-spectral optoacoustic tomography (“MSOT”). MSOT is based on illuminating a tissue with transient laser light and creating pressure variations inside the tissue through a thermo-elastic effect, which leads to acoustic wave propagation. These acoustic waves are conventionally measured at a distance from an inner or outer boundary of the tissue and used to form an image of the energy deposition within the tissue. By using laser light at different wavelengths, a three-dimensional map of tissue constituents and tissue biomarkers can be obtained. This technique has been shown to facilitate the differentiation of various tissue types according to their spectral properties and to image fluorescent probes and nanoparticles that exhibit an absorption resonance in the exciting wavelength. The advantage of this technique over the fluorescence imaging is that it can provide high resolution three-dimensional maps of the concentration of photo-absorbing agents. The ability to localize specific optical agents and tissue constituents in three dimensions enables the differentiation between different probes and tissues in a target region. Additionally, since the anatomy of the imaged tissue and a hollow organ space can be resolved with high resolution, a correction for light attenuation may be performed, leading to the improved quantitative spatial mapping of an agent and concentration of its specific biomarker. Thus, this technique can potentially overcome limitations of fluorescence imaging, namely undesired surface-weighted images and non-quantified results.
Conversely, in comparison to fluorescence imaging, optoacoustic imaging is less sensitive in detecting fluorochromes. This elicits a diagnostic limitation of stand-alone optoacoustic approaches. Although the detection sensitivity may be improved by, for example, including more exciting wavelengths or increasing the signal-to-noise ratio (“SNR”) using averaging, these processes are associated with increased measurement time. While three-dimensional imaging requires hundreds to thousands of slices for proper assessment through visualization, the time required for measurement should ideally not be more than a few minutes for procedures such as intraluminal imaging due to their invasive nature. Furthermore, current approaches for increasing SNR in non-invasive MSOT methods cannot be translated to imaging of hollow organs. For example, non-invasive approaches have been demonstrated for two-dimensional imaging with measurement durations of a few seconds by utilizing multiple detectors and maximizing detector size. However, imaging of hollow organs poses very stringent restrictions on these characteristics. In particular, externally located detectors (that is, on the outside of the vessel) are not feasible for minimally invasive detection, and noninvasive sensors on the outside of the body are unlikely to detect a signal from an intravascular source without severe degradation of the signal. In addition, the sensor size is limited and multiplexing more than one sensor is complex and leads to the reduced SNR per sensor. Some attempts have been made for constructing an intravascular optoacoustic catheter by mounting an intra-vascular ultrasound (“IVUS”) catheter on a thick optical fiber. The sensitivity and speed achieved with this catheter, however, were not sufficient for imaging molecular probes in vivo. Additionally, the total thickness of this two-shaft catheter was a few millimeters, preventing its safe use in human coronary-artery imaging.
Other approaches to improve fluorescence imaging of hollow organs have included the incorporation of optical coherence tomography (“OCT”). Currently, such approaches provide a dual-shaft catheter including a first shaft to perform fluorescence imaging and a second shaft to perform OCT, therefore providing functional information gathered through fluorescence imaging in combination with structural information gathered through OCT. As discussed above, two-shaft catheters require dimensions that are too large for safe use in many intraluminal applications, such as human coronary-artery imaging.
It would therefore be desirable to provide a method and system that is capable of both structural and functional imaging and is also dimensioned to safely perform such imaging for intraluminal applications.